Focused ultrasound (i.e., acoustic waves having a frequency greater than about 20 kiloHertz) can be used to image or therapeutically treat internal body tissues within a patient. For example, ultrasound waves may be used in applications involving ablation of tumors, targeted drug delivery, disruption of the blood-brain barrier (BBB), lysing of clots, and other surgical procedures. During tumor ablation, a piezoceramic transducer is placed externally to the patient, but in close proximity to the tumor to be ablated (i.e., the target region). The transducer converts an electronic drive signal into mechanical vibrations, resulting in the emission of acoustic waves (a process hereinafter referred to as “sonication”). The transducer may be geometrically shaped and positioned along with other such transducers so that the ultrasound energy they emit collectively forms a focused beam at a “focal zone” corresponding to (or within) the target region. Alternatively or additionally, a single transducer may be formed of a plurality of individually driven transducer elements whose phases can each be controlled independently. Such a “phased-array” transducer facilitates steering the focal zone to different locations by adjusting the relative phases among the transducers. As used herein, the term “element” means either an individual transducer in an array or an independently drivable portion of a single transducer. Magnetic resonance imaging (MRI) may be used to visualize the patient and target, and thereby to guide the ultrasound beam.
During a focused ultrasound procedure, a series of sonications is applied to cause coagulation necrosis of the target tissue (such as a tumor) without damaging surrounding tissue. To achieve this, ultrasonic energy emitted from the transducer must be accurately and reliably shaped and focused onto the desired target location. Transducer elements that are not properly configured can lead to improper focal qualities, thereby causing ineffective treatment and/or undesired damage to the non-target tissue. In addition, improperly shaped ultrasound beams may generate unexpected, secondary hot spots at locations other than the intended focal zone; such hot spots may lead to undesired heating, pain for the patient, and/or possibly necrosis of non-targeted tissue.
One source of transducer output errors results from geometric imperfections in the transducer elements (i.e., deviations from their expected locations). For example, assuming a transducer is designed to have a spherical shape, the software that drives each transducer element is configured to activate individual transducer elements based on their positioning according to a spherical model or design. To the extent that the actual location of one or more transducer elements is shifted from the expected location during manufacture, use and/or repair, or if the location shifts as a result of, for example, deformation by heat, the result can be permanent focusing errors due to software programmed according to an ideal spherical model.
Another source of transducer output errors is inhomogeneity of the intervening medium (e.g., a fluid or tissue) through which the ultrasound waves travel prior to reaching the focal zone. The ultrasound waves may interact with the medium through multiple processes, including propagation, scattering, absorption, reflection, and refraction. For example, inhomogeneity of the medium may cause refraction of acoustic energy at the boundaries of regions that have different speeds of sound. Refraction may decrease constructive interference, and hence, the intensity of the acoustic energy at the focal zone. Thus, an inhomogeneous medium may generate beam aberrations and refractions that distort the focus and reduce its intensity, thereby affecting treatment efficiency.
One approach to ameliorating these problems involves focusing the transducer in water at a focal point and using a hydrophone to locate the focal point of maximum intensity. Each transducer element is separately activated at the maximum intensity point, and the phase of each signal is measured by the hydrophone. The measured phase for each element is compared to the expected phase to determine the phase deviation resulting from the geometric imperfections of the transducer elements and/or aberrations resulting from the water; the drive signal is then adjusted to compensate for the observed phase deviation. This approach, however, has a number of shortcomings. For example, because the hydrophone must be placed precisely at the focal point, this point must be identified with precision using, e.g., a highly accurate scanner and electronics; this setup may be expensive. In addition, the transducer elements are tested and calibrated sequentially, which is time consuming. Further, high ultrasonic intensities can damage or even destroy the hydrophone.
Another approach for calibrating the transducer geometric errors and/or beam aberrations resulting from the intervening medium involves placing a point source reflector (e.g., a microbubble) at the focal point. Reflected signals from the point source may be detected and the deviation between the measured phase of the reflected signal and the expected phase (based on the intended focal point) can be determined; the drive signal can then be adjusted to compensate for the deviation. But again, this approach requires an expensive scanner and electronics to identify the focal point so as to align the point source reflector therewith. In addition, at a high acoustic intensity, the point source reflector may produce microbubble cavitation and/or other non-linear effects on the target tissue, which may be difficult to control and which can interfere with the calibration procedure.
Accordingly, there is a need for efficient, economic and reliable approaches to compensating for deviations in the transducer geometry and inhomogeneities in an intervening medium as the ultrasound passes therethrough, thereby creating a high-quality focus.